The present application relates to the art of medical diagnostic imaging in which penetrating X-Radiation is received by an X-Radiation-sensitive detector. The application subject matter finds particular use in computerized tomographic (CT) scanners and will be described with particular reference thereto. However, the invention may also find use in connection with other diagnostic imaging modalities, industrial quality assurance imaging, baggage inspection, X-Ray fluoroscopy, and the like.
Modern X-Ray computer tomography scanners commonly employ X-Radiation detectors to convert X-Ray energy into electrical signals. A detector is usually composed of a scintillator to convert X-Ray energy into light and photosensors, such as a photodiode array, charged-coupled device (CCD) array, etc., to convert that light into an electrical current. The formats of photodiodes used in CT applications include single photodiode elements, one-dimensional (1-D) integrated circuit (IC) photodiode arrays, and two-dimensional (2-D) IC photodiode arrays.
Typically, the electrical signal from each active photodiode element is individually routed to an adjacent pre-amplifier channel. A wire bond connects a top surface bond pad on one end of the photodiode to an external connection. The conductive path to downstream processing electronics is completed using various design options. Pre-amplifiers are either located on the same PC board that includes the detector array or at a more distant location accessed by a cable.
The bond pads are typically located at one end of the photodiodes in sparse 1-D arrays. As the density of elements in the array increases, the bond pads are located on either end of the 1-D array. In some embodiments, the wire bonds in adjacent channels are made at alternate ends.
The wire bond density becomes even more acute for 2-D arrays. A conductive trace from each inner photodiode element in a 2-D array to a connection point for electrical connection with an external connector must be provided. This trace is usually included on the photodiode surface between rows of active photodiode elements. One trace is required per element and each trace usually terminates in a bond pad at an end of the 2-D array. Wire bonds from each bond pad are then made to external connections.
As the number of elements in a 2-D array gets large, two restrictions occur. The space required to provide room for the conductive paths between the detector rows increases and the density of the bond pads at either end of each 2-D array also increases. There is a physical limit, in terms of cost, function, and reliability, as to the number and size of traces and bond pads that can be made using top surface contacts. A conductive path xe2x80x9cbottleneckxe2x80x9d occurs if there is not enough space on a surface to accommodate the number of traces from the photodiode bond pads to the detector electronics.
Also, it is difficult to build a mosaic detector of arbitrary size and shape, i.e., wherein a detector element can abut like detector elements on all sides. Since the contacts for external connections are made laterally, on the sides of the active area, a loss of active area on the receiving surface of the detector results.
Another problem relates to degradation of the signals as they travel over the long bus system between the X-Radiation detectors and the signal processing circuitry.
CT scanners operate in a sea of extraneous radio frequency electromagnetic signals, the frequencies of which vary over a wide band. Sources of extraneous signals include nearby operating electrical components, including high-power tube generators, equipment, signals from other detectors, and the like. The long bus systems include long lead wires which inadvertently act as antennas in picking up extraneous electromagnetic signals and converting them into analog signals. The extraneous analog signals are superimposed on and mix with the analog signals from the detectors. The superimposed extraneous signals appear as noise and fictitious data when reconstructed into images. The resulting images are degraded by noise, ghosting, and other artifacts.
The present invention contemplates an improved X-Radiation detector and CT method and apparatus which overcomes the above-referenced problems and others.
In one aspect of the present invention, a computerized tomography imaging scanner includes an X-Radiation-sensitive detector array for converting received X-Radiation into electrical signals and an image reconstruction processor for reconstructing images based on the received X-Radiation. The detector array includes a scintillation layer converting X-Radiation into visible light and a plurality of back contact photodiode detector modules optically coupled to the scintillation layer, and the detector modules tiled to form a mosaic detector. Each detector module comprises a bounded plane light-sensing surface defining a footprint and an electrical connector for connecting the bounded plane light-sensing surface to readout electronics, each electrical connector being contained within its respective footprint.
In a further aspect, the present invention relates to an imaging system comprising an X-Radiation source selectively generating X-Radiation which at least partially traverses an examination region. The imaging system also comprises an X-Radiation-sensitive layer which converts received X-Radiation into photons of light. Additionally, each X-Radiation detector module includes an array of photodetector devices, each device having a side in optical communication with the X-Radiation-sensitive layer and each generating electrical signals responsive to the photons of light generated by the X-Radiation-sensitive layer. Each detector module includes a carrier substrate supporting each photodetector device, the carrier substrate configured to provide an electrical path from contacts on a back side of the photodetector device through the carrier substrate.
In a further aspect of the present invention, an X-Radiation detector comprises an X-Radiation-sensitive surface which converts received X-Radiation into photons of light, a photosensitive device which has a side in optical communication with the X-Radiation-sensitive surface and which generates electrical signals responsive to the photons of light generated by the X-Radiation-sensitive surface, and a carrier substrate supporting the photosensitive device, the carrier substrate configured to provide an electrical path from contacts on a back side of the photosensitive device through the carrier substrate.
In yet a further aspect, the present invention relates to an imaging method comprising irradiating an X-Radiation-sensitive surface with X-Radiation, converting the X-Radiation incident upon the X-Radiation-sensitive surface into light, transmitting the light to a light-sensitive surface of a photodiode array to produce an electrical signal proportional to the converted light, and communicating the electrical signal to signal processing circuitry via a conductive path. The conductive path comprises a first set of contacts disposed on a back surface of the photodiode array opposite the light-sensitive surface; a second set of contacts disposed on a front surface of a circuit carrying substrate, the second set of contacts aligned with and electrically coupled to the first set of contacts when the back surface of the photodiode array and the front surface of the substrate are in aligned facing relation; and a third set of contacts disposed on a back surface of the substrate opposite the front surface of the substrate.
In another aspect, a front-illuminated photodiode array is provided, in which signals from the photodiode elements are passed to the photodiodes backside using conductive vias. The vias are made conductive by impurity diffusion of the substrate silicon. On the backside, a metal pad forms an ohmic contact to the diffused region of the via. Bump bonds are then made to these backside contacts. The impurity diffused via forms a reverse bias diode contact with respect to the subsrate silicon and thus an independent conductor for the photodiode signal from the topside.
One advantage of the present invention resides in locating the electrical conductors from the photodiode beneath the photodiode array.
Another advantage in one embodiment of the invention resides in freeing the light-sensitive surface from electrical conductors.
Another advantage of the present invention resides in the ability to disperse a plurality of electrical leads or traces through a multi-level substrate.
Another advantage in one embodiment of the invention is that the present invention increases the active surface area of the photodiodes available to receive X-Rays.
Another advantage is that it improves X-Ray conversion efficiency.
Another advantage of the present invention resides in the ability to group a plurality of detector arrays together into a variety of configurations.
Yet another advantage of the present invention is that signal processing electronics can readily be shielded from X-Radiation.
Another advantage resides in the ability to reduce the number of contacts by incorporating multiplexing circuitry as a part of the functional integrated circuitry of the photodiode array.
Still further advantages will become apparent to those of ordinary skill in the art upon reading and understanding the following detailed description.